The present embodiments relate to a radiation detector that may be used to detect electromagnetic radiation.
Imaging systems appertaining to medical technology are becoming increasingly important nowadays. Systems of this type are used to generate two- or three-dimensional image data of organs and structures of the human body, which may be used, for example, for diagnosing causes of illness, for carrying out operations, and for preparing therapeutic measures. The image data may be generated based on measurement signals obtained with the aid of a radiation detector.
This is the case, for example, in X-ray and computed tomography systems (CT). In systems of this type, the body or a body section of a patient to be examined is radiographed by X-ray radiation generated by a radiation source. The non-absorbed, transmitted portion of radiation is detected by a detector.
A further example is image generation with the aid of radionuclides, such as is used in positron emission tomography systems (PET) and single photon emission computer tomography systems (SPECT). In this case, the patient to be examined is injected with a radiopharmaceutical that generates gamma quanta either directly (SPECT) or indirectly (PET) through emission of positrons. The gamma radiation is detected by a corresponding radiation detector.
Detectors that may be used for the energy-resolved detection or “counting” of radiation quanta may operate according to different measurement principles. Radiation may be detected either directly (e.g., by direct conversion of the radiation energy into electrical energy) or indirectly. In the case of the last-mentioned variant, use is generally made of a scintillator, which is excited in response to the action of radiation to be detected and reemits the excitation energy by emitting lower-energy electromagnetic radiation. Only the radiation emitted by the scintillator is converted into electrical measurement signals in this case. Detectors of planar construction (e.g., “flat detectors”) that are used in the medical field and operate in accordance with these measurement principles are described, for example, in M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol (2005), 15: 1934-1947.
The conversion of the radiation emerging from a scintillator into an electrical signal may be effected in various ways. Besides use of a photomultiplier provided with a photocathode in the form of an evacuated electron tube, one concept that is common at the present time includes using a silicon photomultiplier (“SiPM”). This involves a matrix arrangement of avalanche photodiodes (APD) embodied on a shared substrate, electrons being generated in the photodiodes as a result of incident photons, and the electrons being multiplied in an avalanche-like manner.
One disadvantage of silicon photomultipliers, however, is that only part of the total area available for irradiation may be utilized as sensitive or “active” area. The reason for this is that between the active or radiation-sensitive regions there are also insensitive regions, in which resistors and signal lines or wiring structures are arranged. A silicon photomultiplier therefore has a relatively small ratio of active area to total area (e.g., irradiated total area). The ratio is also designated as “filling factor.” Further disadvantages include noise that occurs during operation, and a relatively high dark rate or dark count. In other words, signal generation takes place even without irradiation.
A detector including a scintillator and a silicon photomultiplier may be embodied such that the silicon photomultiplier is opposite an end face or rear side of the scintillator. An opposite end face or front side of the scintillator faces the radiation to be detected. As a result, the silicon photomultiplier may detect only that portion of the radiation converted in the scintillator that emerges at the rear side thereof. Proceeding from the respective excitation or interaction location in the scintillator, however, the scintillation radiation is emitted not only in the direction of the rear side, but also in other directions. The radiation is subject to loss processes such as reflection, absorption and scattering. In the case of scintillators having a high aspect ratio (e.g., a high ratio of height to width), as may be the case for example in a PET system, the losses are therefore relatively high. In the case of an aspect ratio of greater than 7:1, the radiation emerging from a scintillator may make up a proportion of merely 40-60% of the total radiation generated. Although a higher intensity of the incident radiation may be provided in order to compensate for the losses, as a result, a patient is also exposed to an increased radiation dose.
It is disadvantageous that an interaction location of incident radiation in the scintillator may not be detected or may be detected only with very great difficulty based on the radiation emerging at the rear side of the scintillator. It is not possible to obtain information about the height or depth of an interaction in the scintillator. Such disadvantages therefore restrict the resolution of an imaging system provided with such a detector construction.
For image intensification and for electron multiplication, it is furthermore known to use microchannel plates (MCP) having a number of channels. During operation, an electrical voltage present along the channels is generated, whereby entering electrons may be accelerated within the channels and multiplied by impacts with the channel walls. Use of a microchannel plate in connection with an image intensifier is described in US 2009/0256063 A1, for example.